This invention relates generally to the field of magnetic resonance imaging (MRI). Specifically, the invention relates to spin-locking and T1xcfx81-weighted MRI pulse sequences.
Magnetic resonance imaging (MRI) has become the modality of choice for imaging joints due to its excellent definition of ligaments, cartilage, bone, muscle, fat and superior soft tissue contrast (Smith, Magn. Reson. Imaging Clin. N. Am. 3:229-248 (1995); Sofka, et al., Radiology 5:217-226 (2001)). For two decades, proton magnetic resonance imaging (MRI) has shown its efficacy in the noninvasive analysis of soft tissues, particularly in the diagnosis of tendinomuscular and osteoarticular diseases (Peterfy, et al., Radiol. Clin. North Am. 34:195 (1996); Peterfy, Magn. Reson, Imaging Clin. N. Amer. 8:409-430 (2000)).
Articular cartilage is a connective tissue consisting of relatively few cells and a highly charged and hydrated extracellular matrix (ECM). The constituents of the ECM are proteoglycans (PG), collagen, and non-collagenous proteins and water (Grushko, et al., Conn. Tiss. Res. 19:149-176 (1989); Lohmander, J. Anatomy 184:477-492 (1994); Mankin, et al., J. Bone Joint Surg-Am. 53:523-537 (1971)). Despite its remarkable durability, degeneration of articular cartilage can result from either noninflammatory processes, such as osteoarthritis (OA), or inflammatory processes, such as rheumatoid arthritis (RA). The early stage of OA is associated with loss of PG and changes in water content (Grushko, et al., 1989; Lohmander, 1994). Recent developments in chondroprotective therapies, cartilage grafting, gene therapy and tissue engineering have increased the demand for accurate and non-invasive techniques that will enable the detection of the early biochemical changes of cartilage degeneration in vivo.
Conventional proton MR techniques have been able to provide information about late stages of degeneration in which structural defects are present (Recht, et al., Am. J. Roent. 163:283-290 (1994); Peterfy, et al., Radiol. Clin. North Am. 32:291-311 (1994)). Recently, delayed gadolinium (Gd)-enhanced proton MRI of cartilage (dGEMRIC) (Bashir, et al., Magn. Reson. Med. 36:665-673 (1996); Burstein, et al., Magn. Reson. Med. 45:36-41 (2001); Mlynarik, et al., J. Magn. Reson. Imaging 10:497-502 (1999)), positively charged nitroxide based techniques (Lattanzio, et al., 25:1024 (2000), and sodium MRI (Reddy, et al., Magn. Reson. Med. 39:697-701 (1998); Shapiro, et al., J. Magn. Reson. 142:24-31; Shapiro, et al., Magn. Reson. Med. 47:284-291 (2002)) have been employed to measure PG changes in cartilage both in vivo and in vitro. However, these techniques have some practical limitations. In dGEMRIC, long waiting period between contrast agent injection and imaging and variation in intra tissue Gd-relaxivity may contribute to the errors in PG quantitation, thereby reducing the accuracy of this technique in the detection of OA. Although sodium MR imaging has high specificity towards proteoglycans, it has an inherently low sensitivity and requires special radio-frequency hardware modifications before it can be used with a routine clinical imaging unit.
Spin lattice relaxation time in the rotating frame (T1xcfx81) has been shown to be sensitive to changes in PG content of cartilage (Duvvuri, et al., Magn.Reson.Med. 38:863-867 (1997); Akella, et al., Magn. Reson. Med. 46:419-423 (2001)). It is well suited for probing macromolecular slow motions at high static fields without modifying MR system hardware (Sepponen, et al, J, Computer Assisted Tomography 9:1007-1011 (1985); Santyr, et al., Magn. Reson. Med. 12:25-37 (1989)).
T1xcfx81 provides an alternative contrast compared to conventional MRI methods. Since the first description by Redfield (Phys. Rev. 98:1787 (1955)), spin-locking technique has been used extensively, to investigate the low frequency interactions between the macromolecules and bulk water. Several authors have investigated the T1xcfx81 dispersion characteristics of biological tissues, including brain (Aronen, et al., Magn. Reson. Imag. 17:1001-1010 (1999); Rizi, et al., J Magn Reson Imaging 8:1090-1096 (1998)), tumors (Aronen, et al., Magn. Reson. Imag. 17:1001-1010 (1999); Markkola, et al., Radiology 200:369-375 (1996)), and articular cartilage (Mlynarik, et al., 1999; Akella, et al., 2001; Duvvuri, et al., 1997; Duvvuri, et al., Radiology 220:822-826 (2001)). These studies have demonstrated the potential value of T1xcfx81-weighting in evaluating various physiologic/pathologic states.
Although recent studies have demonstrated the potential role of T1xcfx81-weighted MRI in measuring cartilage degeneration, they all have been restricted to single slice imaging, and hence, are impractical for the imaging of a typical anatomic volume. The use of single slice techniques results from the problem in making the spin-locking pulse slice selective. A 2D multi-slice T1xcfx81 MRI sequence has not been implemented since the use of a nonselective spin-lock (SL) pulse poses a challenge to slice-selective imaging. Multi-slice imaging requires the application of multiple radio frequency (RF) pulse trains within a sequence repetition time (TR) to excite several slices in a time efficient manner. Currently, T1xcfx81 pulse sequences employ a non-selective RF pulse to spin-lock the magnetization in the transverse plane following the application of a non-selective xcfx80/2 pulse. As a result, using this method, the spin-lock pulse excites signals from the entire sample during each application, but the subsequent imaging sequence acquires data from only a single slice, wasting the information from the remainder of the volume. Accordingly, a need has remained until the present invention for a method to perform multi-slice T1xcfx81-weighted MRI.
Furthermore, in addition to the improved ability to image anatomic regions through 2D multi slice imaging, a 3D image provides even better contrast and faster acquisition than a 2D image. Even though a 3D, gradient-echo readout of a T1xcfx81-weighted MR signal has been used (Aronen, et al., 1999), that sequence was implemented on a low field magnet (0.1T) with a combination of adiabatic pulses, and RF spoiling alone was employed to destroy unwanted transverse coherence. The use of adiabatic pulses has certain drawbacks. Their long pulse lengths result in substantial decay of magnetization during the pulse period. These pulses cannot be easily calibrated on a clinical scanner, are more RF power intensive and may introduce specific absorption rate (SAR) issues. Moreover, any residual transverse magnetization resulting from incomplete restoration of the T1xcfx81-prepared magnetization to the longitudinal axis by the second adiabatic pulse will result in unwanted image artifacts.
The need has remained for a MR pulse sequence capable of performing 3D T1xcfx81-weighted MRI imaging on a high field clinical scanner. However, when using high field clinical scanners (greater than or equal to 1.5T), the SAR by the pulse sequence is significant, and imaging parameters must be chosen such that the energy deposition does not exceed the established SAR guidelines. Hence, determining the optimal sequence parameters has been necessary, so that the energy deposition by the radio frequency pulses in the sequence, measured as the SAR, does not exceed safety guidelines for imaging human subjects.
Additionally, significant artifacts arise in T1xcfx81-weighted imaging when nutation angles suffer small deviations from their expected values. These artifacts vary with spin-locking time and amplitude, severely limiting attempts to perform quantitative imaging or measurement of T1xcfx81 relaxation times. As a result, a need has also remained for a xe2x80x9cself-compensatingxe2x80x9d spin-locking pulse that dramatically reduces artifacts and provides a more robust implementation of T1xcfx81 imaging despite spatial variations in nutation angles.
Finally, T1xcfx81-weighted MRI in volume coils has been employed to improve tissue contrast (Markkola, et al., J. Magn. Reson. Imaging 7:873-879 (1997); Santyr, Magn. Reson. Imaging Clin. N. Am. 2:673-690 (1994)). However, the use of surface coils is attractive due to their signal-to-noise ratio (SNR) characteristics, especially when used on extremities. Inhomogeneous B1 profiles of surface coils produce undesired image artifacts that prevent conventional T1xcfx81-weighted imaging. Accordingly, a need has remained for combining 3D T1xcfx81-weighted MRI and a self compensating spin-lock pulse to obtain artifact-free 3D T1xcfx81-weighted MRI using surface coils.
The present invention provides novel T1xcfx81-weighted imaging pulse sequences for 2D multi-slice and 3D T1xcfx81 MR imaging. Also provided is a pulse sequence for correction of artifacts with a self-compensating spin-locking pulse.
In one aspect of the present invention, a 2D multi slice T1xcfx81-weighted MR imaging pulse sequence is provided. The pulse sequence comprises a first xcfx80/2 pulse for selectively exciting a band of spins that are spin-locked in the transverse plane by the application of a train of spin-locking pulses with alternating phase (xc2x190xc2x0 phase-shifted from the phase of the first xcfx80/2 pulse), a second slice-selective xcfx80/2 pulse which restores the spin-locked magnetization to the longitudinal axis for imaging with any pulse sequence, and a strong xe2x80x9ccrusherxe2x80x9d gradient applied to destroy any residual transverse magnetization. Also provided is a method for obtaining 2D multi-slice T1xcfx81-weighted MR images
In another aspect of the invention, a self-compensating spin-locking pulse sequence for correcting artifacts in T1xcfx81-weighted imaging is provided. A preferred embodiment of the self-compensating spin-locking pulse comprises a four-pulse preparatory cluster, further comprising a xcfx80/2 pulse (P1) which nutates magnetization into the transverse plane, followed by a spin-locking pulse (SL) with phase-shift of 90xc2x0 and amplitude Bs1, which is applied for some time TSL/2, then an SL pulse with phase-shift of xe2x88x9290xc2x0 and amplitude Bs1, and a final xcfx80/2 pulse (P2) which is applied, phase-shifted by 180xc2x0 to return the T1xcfx81-relaxed magnetization to the longitudinal axis for imaging. Bs1 is some fraction of the maximum available RF field B1. A crusher gradient is then applied to dephase any residual longitudinal magnetization. Additionally, a method of correcting for artifacts in T1xcfx81-weighted MR imaging is provided.
In yet another aspect of the invention there is provided a 3D T1xcfx81-weighted pulse sequence for MR imaging. A preferred embodiment of the 3D T1xcfx81-weighted pulse sequence comprises: a first xcfx80/2 hard pulse; a spin-lock pulse having amplitude B1; a second xcfx80/2 hard pulse phase-shifted 180 degrees from the first hard pulse; and a strong xe2x80x9ccrusherxe2x80x9d gradient applied before the xcex1 pulse to destroy any residual magnetization in the transverse plane and prevent the formation of unwanted coherences. A method of obtaining 3D T1xcfx81-weighted MR images is also provided. In yet another embodiment a method of obtaining 3D T1xcfx81-weighted MR images with a surface coil is provided.
Additional objects, advantages and novel features of the invention will be set forth in part in the description, examples and figures which follow, and in part will become apparent to those skilled in the art on examination of the following, or may be learned by practice of the invention.